Electro-diffusion enhanced bio-molecule charge detection using electrostatic interaction

ABSTRACT

According to one aspect, the disclosure is directed to an example embodiment in which a circuit-based arrangement includes a circuit-based substrate securing a channel, with an effective width that is not limited by the Debye screening length, along a surface of the substrate. A pair of reservoirs are included in or on the substrate and configured for containing and presenting a sample having bio-molecules for delivery in the channel. A pair of electrodes electrically couple a charge in the sample to enhance ionic current flow therein (e.g., to overcome the electrolyte screening), and a sense electrode is located along the channel for sensing a characteristic of the biological sample by using the electrostatic interaction between the enhanced ionic current flow of the sample and the sense electrode. Actual detection occurs by using a charge-signal processing circuit to process the sensed charge signal and, therefrom, provide an output indicative of a signature for the bio-molecules delivered in the channel.

RELATED PATENT DOCUMENTS

This patent document claims the benefit, under 35 U.S.C. §119(e), ofU.S. Provisional Patent Application Ser. No. 61/220,496 filed on Jun.25, 2009, and entitled “Electro-Diffusion Enhanced Bio-Molecule ChargeDetection Using Long-Range Electrostatic Interaction;” this patentdocument and the Appendices A-D filed in the underlying provisionalapplication are fully incorporated herein by reference.

FIELD

The present disclosure is generally directed to methods and apparatuses,including but not limited to semiconductors, providing for and usingelectrical detection of bio-molecules charges.

BACKGROUND

The electrical detection of bio-molecules is currently an activeresearch field. Two types of detection approaches are typically used.The first type is based on translocation of bio-molecules throughnanometer-size ion channels across biological membranes or solid-statepores across dielectric membranes. The detection mechanism is based onthe blockage of the electrical conductance of the pores duringtranslocation events. This approach is capable of single bio-moleculeanalysis. However, due to the small sizes of the bio-molecules (DNAs inthe referred studies), the size of the pores is required to be small(<˜10 nm) for appreciable current blockage signals, which imposesenormous fabrication challenges. The second type uses field effect todetect the biological charge of bio-molecules that are specificallyimmobilized to the sensor surface; this includes both planar ISFETstructure and, more recently, nanowire structures. This affinity-typeapproach often detects the ensemble average signal of the immobilizedbio-molecules. Electrolyte screening is known to impose a fundamentallimit on its charge sensing capabilities at physiological condition. Asan example, the characteristic Debye screening length in a physiologicalcondition (˜100 mM salt concentration) is ˜1 nm.

In other research efforts, scientists have described DNA sequencingdevices that use semiconductor field effect transistors at the surfaceof an opening or recess for charge detection of analytes flowing by inthe close vicinity under electrophoresis process. Just as the ISFET typesensing scheme, such proposed devices are still limited by theelectrolyte screening since no mechanisms are identified or designed toovercome that limit. Their charge sensing operation is dependent on theclose proximity (˜Debye length) of bio-molecules to the sensingelements. Such devices propose relatively small-size openings (e.g.,<˜10 nm), which enormously complicates the device fabrication steps.These devices also use metal and insulator layers to replace the p-dopedor n-doped semiconductor regions. The operation of this type of deviceis based on the principle that the field effect will change thetunneling current across the insulator layer from one metal layer toanother.

Accordingly, there is need for apparatuses and methods, involving thedetection of bio-molecule charges, that overcome these and otherlimitations.

SUMMARY

Certain aspects of the present disclosure are directed to apparatusesand methods that overcome the above-mentioned and other issues.Consistent therewith, specific apparatuses and methods involve thedetection of bio-molecule charges based on their long-rangeelectrostatic interaction. In one of various embodiments, such detectionis directed to bio-molecule charges with significantly-enhancedsensitivity by exploiting electro-diffusion ionic current flow inmicro-machined channels and other types of channels (e.g., laterally-and vertically-oriented).

According to a specific example, the present disclosure is directed toimplementations of an apparatus and/or method involving the enhanceddetection of charges of bio-molecules traversing a channel situatedadjacent a sense electrode and between biasing electrodes; such chargedetection is not limited by the electrolyte screening. In more specificembodiments of the disclosure, aspects include (operating alone or incombinations thereof): the channel being largely compatible withconventional manufacturing requirements (e.g., in terms of foundry andmicro-electromechanical system processes and/or feature size); thebio-molecule charges being sensed using a long-range electrostaticinteraction; the electrodes being controlled to manipulate the charges,the charge screening and/or movement of the bio-molecules forfacilitating detection; the inclusion of AC modulation and lock-intechniques for enhanced sensing applications; and the detection ofbio-molecule charges at the single-molecule level.

In connection with other specific embodiments, the disclosure isdirected to a system for detecting bio-molecule charges based on theirelectrostatic interaction. The system includes an integrated-circuitarrangement that detects bio-molecule charges, the arrangement includesa circuit-supporting substrate, at least one reservoir configured forcontaining and presenting a sample having bio-molecules, first andsecond electrodes configured for electrically coupling a charge in thesample, at least one sense electrode located along the channel, and acharge-signal processing circuit. The circuit-supporting substrateincludes a channel that is not limited by the Debye screening length. Inone implementation, the channel is located along a surface of thesubstrate (e.g., micro-machined along the surface or formed via anattached separate member) and, in another implementation, the channel isformed using a through hole in the substrate. The reservoir is alsoconfigured for delivering the bio-molecules in the channel, and thebiasing electrodes are configured for enhanced ionic current flow, e.g.,overcoming the electrolyte screening, and therefore enhancing theelectrically coupling a charge in the sample to the sense electrodes. Inresponse to the sense electrode detecting a characteristic of thebiological sample based on the electrostatic interaction with thesample, the charge-signal processing circuit communicatively provides anoutput indicative of a signature for the bio-molecules delivered in thechannel. In a more specific implementation, the integrated-circuitarrangement has at least one of the electrodes buried in the substrateadjacent the channel and used in this configuration for charge sensingand/or bio-molecule movement control. In another more specificimplementation, the integrated-circuit arrangement employsnanostructures (e.g., nanowires or nanotube arrays) along the surfacechannel, e.g., along the bottom of a laterally-arranged channel, forsensing charges of the translocating biomolecules.

Other embodiments are indicated by way of the examples described herein,and by way of the contemplated claims and papers appended hereto.

The above summary is limited to characterizing certain aspects and isnot intended to describe each illustrated embodiment or everyimplementation of the present disclosure. The figures and detaileddescription that follow, including that described in the appendedclaims, more particularly exemplify embodiments in support of thedisclosure.

BRIEF DESCRIPTION OF THE DRAWINGS

The present disclosure may be more completely understood inconsideration of the detailed description of various embodiments of thedisclosure that follows in connection with the accompanying drawings asfollows:

FIGS. 1 a, 1 b and 1 c are illustrations of exemplary biosensorapparatuses consistent with the present disclosure, with FIGS. 1 a and 1b respectively providing a plan view and a perspective view of anexemplary biosensor apparatus and FIG. 1 c showing a plan view of analternative implementation for sensing electrodes in another exemplarybiosensor apparatus.

FIGS. 2 a and 2 b are illustrations of alternative biosensor sensordevices, according to the present disclosure.

FIG. 3 is a cross-sectional view of an exemplary biosensor apparatus,according to the present disclosure.

FIG. 4 is a cross-sectional view of an exemplary biosensor apparatusincluding a through-substrate (or slab) channel implementation,according to the present disclosure.

Also consistent with embodiments of the present disclosure,

FIGS. 5 a, 5 b and 5 c are diagrams useful in showing use of charge tagsand antibody/antigen reaction to achieve high specificity; and

FIG. 6 a is a block diagram and FIGS. 6 b, 6 c and 6 d are graphs usefulin showing the effect of proper electrical biasing in embodimentsconsistent with the present disclosure.

While the disclosure is amenable to various modifications andalternative forms, specifics thereof have been shown by way of examplein the drawings and will be described in detail. It should beunderstood, however, that the intention is not to limit the disclosureto the particular embodiments described. On the contrary, the intentionis to cover all modifications, equivalents, and alternatives fallingwithin the spirit and scope of the disclosure.

DETAILED DESCRIPTION

The present disclosure is believed to be useful for a variety ofapplications involving the sensing/detection of bio-molecules from theircharacteristic charges. As nonlimiting examples, these applicationsinclude gene sequence applications such as those involving the detectionof DNA hybridization based on oligomer charges doubling afterhybridization, applications involving immune responses based on changeof charges from antibody-antigen binding, detection of proteinphosphorylation as in applications involving neutral hydroxyl groupreplaced by negatively charged phosphate and other applications forgeneral charge-based particle detection such as charged particles inindustrial waste. While the present disclosure is not limited to suchapplications, various aspects of the disclosure may be appreciatedthrough a discussion of various examples using this context.

Aspects of the present disclosure are directed to the sensing anddetection of bio-molecules in a sample by bio-molecule charges withsignificantly enhanced sensitivity. By exploiting electro-diffusionionic current flow channels without being limited by the Debye screeninglength, embodiments involve use of electrodes for electrically couplinga charge to and for sensing, based on long-range electrostaticinteraction, a characteristic of the sample. Biological macromoleculessuch as DNAs and proteins carry a certain amount of charges in aphysiological electrolyte solution, and the electrical sensing of suchbiological charges is used to gain access to direct detection andanalysis of biological species, ultimately at a single-molecule level.

Certain embodiments of the present disclosure concern the Debye-Huckeltheory which dictates that charge sensing is only possible within a fewDebye lengths from the biomolecules (one Debye length is ˜1 nm at 100 mMelectrolyte concentration that approximates the physiologicalcondition); this theory suggests a detection-sensitivity limit ofelectrostatic screening by electrolyte counter-ions. Consistent withcertain implementations, such a limit can be overcome by carrying outcharge detection under an electrical current flow environment inchannels having an effective width that is not limited by the Debyescreening length; these channels include, e.g., lateral channels such asmicro-machined surface channels and vertical channels as in substratevias. Recently-published surprising simulation results made inconnection with the present disclosure establish that the presence ofsteady-state electro-diffusion ionic flow can dramatically suppress theelectrolyte screening and therefore improve the signal levels. Certainaspects consistent with this approach are based on long-range electricalsensing to significantly relax fabrication complexity and to providecompatibility with integrated circuit (IC) processes. In this regard,these aspects of the present disclosure are highly suitable forlow-cost, point-of-care medical diagnostic applications. For furtherinformation regarding charge detection of bio-molecular substances andthese surprising simulation results, reference may be made to, Y. Liu,K. Lilja, C. Heitzinger, and R. Dutton, Overcoming the Screening-InducedPerformance Limits of Nanowire Biosensors: A Simulation Study on theEffect of Electro-Diffusion Flow, Electron Devices Meeting, IEDM 2008,IEEE International, pp. 1-4 (2008); this document forms part of thispatent document and is attached to the above-noted provisional patentdocument as Appendix A.

Turning now to the drawings, FIGS. 1 a and 1 b are respective plan andperspective views of an exemplary biosensor apparatus according to thepresent disclosure. Channel 110 is depicted between (delivery) reservoir112 a and (receiving) reservoir 112 b, with fluidic interfaces 114 a and114 b connecting sample fluid between the channel 110 and the associatedreservoir 112 a/112 b. A charge-sense electrode 120 is buried near thebottom of the channel 110 and configured to sense charges on theanalytes (charged biomolecules) in response to voltage(s) applied tobiasing electrodes 124 a-b by a voltage bias circuit (not illustrated).The sense electrode 120 is connected to a read-out circuit (notillustrated) that includes charge sensitive amplifiers and, following,stages of signal-multiplexing and signal-processing units. Byimplementing the channel 110 with a width that is not limited by theDebye screening length, an electrolyte solution is introduced to fillthe channel via (delivery) reservoir 112 a and fluidic interface 114 a,a voltage bias circuit controls the biasing electrodes 124 a-b (placedin the solutions in each of the reservoirs), and the analytes areintroduced in one of the chambers and are subjected to an electric fieldensuing from the voltage being applied between (or across) the twobiasing electrodes. The biomolecules travel across the channel to theother side of the channel by electrophoresis. When the biomoleculestravel by the sense electrodes, image charges are induced in theelectrode by field effect, which leads to signals in the read-outcircuits. The multiplexing of signals from the multiple sense electrodesare used to enhance the signal levels and extract information on thetravelling speed of the analytes. As discussed in connection with theincluded appendices, the experimental studies and embodimentsdemonstrate remarkable and unexpected selectivity and sensitivity.

An important option for the above-illustrated detector arrangementsinvolves use of modulation circuit for driving modulation electrodes 128a and 128 b. This modulation is used in combination with a signallock-in circuit to detect and lock onto the correct (modulated) signalas the biomolecules travel across the channel to the other side of thechannel by electrophoresis. When the biomolecules travel by the senseelectrodes, image charges are induced in the electrode by field effect,which leads to discernment of the modulated signals in the read-outcircuits. This discernment is provided for any of various signal lock-incircuits including, for example, largely conventionally-implementphase-locked or frequency-locked loop circuits.

FIG. 1 c illustrates an alternative embodiment in which a charge-senseelectrode 130 (as opposed to the charge-sense electrode 120 of FIG. 1 b)is a nanostructure (such as a nanowire, carbon nanotube or otherreactive nanostructure) which may be functionalized to provide adetectable conductance-type reaction as the biomolecules travel throughthe channel by electrophoresis. In this regard, the sense electrode 130acts like a resistor that changes resistance in the presence of acertain type of complementary molecule.

For an example of work performed in this area of nanostructurefunctionalization, reference may be made to the articles: W. Huang, S.Taylor, K. Fu, Y. Lin, D. Zhang, T. W. Hanks, Apparao M. Rao, andYa-Ping Sun, “Attaching Proteins to Carbon Nanotubes viaDiimide-Activated Amidation,” American Chemical Society, Nano Letters, 2(4), pp. 311-314 (Mar. 16, 2002), and M. Shim, N. Wong Shi Kam, R. J.Chen, Y. Li, and H. Dai, “Functionalization of Carbon Nanotubes forBiocompatibility and Biomolecular Recognition,” American ChemicalSociety, Nano Letters (Jan. 21, 2002).

FIGS. 2A and 2B are illustrations of alternative biosensor sensordevices, according to the present disclosure. FIG. 2A is a simplifiedblock diagram (from a top view) of a circuit-based biosensor sensordevice having a peripheral lock-in amplification circuit 210 with a line212 for the feedback signal used for the lock-in technique appropriatefor the modulation applied at control electrodes 228, as would beappreciated by the skilled artisan. Rather than using a singlecharge-sense electrode, the biosensor sensor device of FIG. 2A uses apair of electrodes 228 but otherwise similar voltage biasing viaelectrodes 220.

FIG. 2B provides a front view of the substrate (or slab) 240 of anotherbiosensor sensor device, according to the present disclosure, in whichthe channel 250 traverses over the (dielectrically-insulated) senseelectrodes. The skilled artisan would appreciate that aspects ofembodiments disclosed herein need not be limited to conventionsemiconductive substrates; thus, while the slab 240 is illustrated as asubstrate-like material having an insulating-oxide, the term “slab” is amore generic term that includes such specific substrates as well asother forms and materials (e.g., plastic) which can be used to supportsuch structures as the channels, electrodes, etc.

In alternative or complementary embodiments (e.g., similarly describedin Appendix B of the above-noted provisional document), the signalsprovided from the sense electrodes 228 can be passed to respectivelyassigned (optionally low-noise) amplifiers for processing the signalpaths along dedicated channels via a signal process circuit. In one morespecific embodiment, the signal process circuit receives the outputsignals via the amplifiers along such dedicated channels and processesthe signals carried thereon by sampling and/or multiplexing these inputsignals.

Using this approach, sensor devices can be implemented in various forms.As just one example, such a sensor device employs a physical sensorchannel stemming from a buried electrode to detect the charge ofbio-molecules that are electrically driven through the surface channels.

For one such (experimental) sensor device, a surface channel (a widthand height ˜100 nm and length ˜1 micrometer) is fabricated usingMicro-Electro-Mechanical Systems (MEMS) technologies and connected withtwo chambers that serve as reservoirs. An array of sense electrodes areburied at the bottom of the channels; they are then connected toread-out circuits that consist of charge sensitive amplifiers andfollowing stages of signal multiplexing and processing units.Electrolyte solutions are introduced to fill in the channel and bothreservoirs. Two biasing electrodes are placed in the solutions in eachof the reservoirs. The analytes (charged biomolecules) are thenintroduced in one of the chambers. By applying voltages between the twobiasing electrodes, the biomolecules travel across the channel to theother side by electrophoresis. When the biomolecules travel by the senseelectrodes, image charges are induced in the electrode by field effect,which leads to signals in the read-out circuits. The multiplexing ofsignals from the multiple sense electrodes can be used to enhance thesignal levels; and extract information based on travelling speeds of theanalytes.

With the surface channel configuration using metal as sensing electrodesand the integrated read-out circuits being highly compatible withfoundry IC and MEMS processes, this approach can lead to improved yieldand reduced cost. Such long distance charge sensing is possible becausethe electrolyte screening is greatly suppressed in the presence of theionic current flow. This scheme senses charge of each analyte moleculethat travels across the sense electrode, i.e., its detection is at asingle biomolecule level. This enables detection with very small samplevolume, which can be advantageous compared to affinity type sensingschemes, whose signal is from an ensemble average of multiple analytemolecules attached to the sensor surface. Also multiple sense electrodescan be multiplexed to enhance signal levels using spatially correlatedsampling techniques. Furthermore, based on the long range electrostaticinteraction, the metal electrodes can be employed as actuators tocontrol the movement of the bio-molecules. It has also been discoveredherewith that the screening suppression effect is a function of theexternal electric field, which means the sensed signal can beelectrically modulated. Based on this finding, AC modulation and lock-intechniques can be employed to dramatically enhance the sensitivities.

The operation principle of this type of charge sensor has been examinedand discussed in Y. Liu, J. Sauer, and R. W. Dutton, “Effect ofelectro-diffusion current flow on electrostatic screening in aqueouspores,” Journal of Applied Physics, vol. 103, p. 084701 (2008), wherethe charge sensing in a relatively large channel is feasible because ofthe electro-diffusion induced screening suppression effect. Underlyingthis effect is the Debye screening equilibrium being disrupted in thepresence of ionic current flow. Using the same numerical technique, afeasibility study of a prototypical surface channel sensor is given inthe following. Using a prototypical 3D device structure as in FIGS. 2 aand/or 2 b, an electrical bias V is applied between the cathode andanode to introduce the electro-diffusion flow. The bio-molecule ismodeled as a water impermeable rod with its dimensions and chargescorresponding to those of a 65-basepair dsDNA fragment. The iontransport is modeled by the coupled Poisson-Nernst-Planck equations. Thesense electrode is modeled as a virtual ground. For various longitudinallocations of the bio-molecule, the percentage of the induced imagecharge at the sense electrode with respect to the total bio-molecularcharge is computed. Along the vertical direction, the bio-molecule isplaced 51 nm away from the sense electrode. The electrolyteconcentration used in the simulations is 1 mM, corresponding to a Debyelength of 10 nm.

Another possible variation is to use Si-nanowire (NW) or carbon nanotube(CNT) arrays as the sensing elements in replacement of the metal sensingelectrodes (FIG. 5). In this scheme, the input charge signal istransduced and amplified to NW/CNT current signals right at the frontend. Therefore, the parasitic effects are minimized: both signal levelsand noise immunity can be greatly improved.

Yet another possible variation is to use affinity-type operation insteadof the translocation-type. As in the common approach of affinity-basedbiosensors, the surface of the sensing elements (metal electrodes,Si-NWs or CNTs) can be chemically functionalized to immobilize analytes.In the common approach, there is only one reference electrode used inthe electrolyte solution. The improvement to the approach is tointentionally introduce the ionic current flow to greatly enhance thesignal levels; as studied in Y. Liu, K. Kilja, C. Heitzinger, and R. W.Dutton, “Overcoming the screening-induced performance limits of nanowirebiosensors: a simulation study on the effect of electro-diffusion flow,”in IEDM Tech. Dig., San Francisco, pp. 491-494 (2008), the signal levelcan by enhanced by about an order of magnitude with ˜100 kV/cm electricfield in the solution. Furthermore, the AC modulation and lock-intechniques can be applied in this design as well.

FIG. 3 is a cross-sectional view of yet another exemplary biosensorapparatus, according to the present disclosure. As with the embodimentsshown in FIGS. 1 a-1 c, bias electrodes 324 are located near or underthe end of the channel 310, and modulation/control electrodes 328 areused when a lock-in amplification circuit (e.g., 210 of FIG. 2 a) isused for enhancing detection capabilities. As illustrated in FIG. 3, thestructures providing the reservoirs 312 and channel 310 at least in partimplemented by different layers which can be built by various techniqueswhich range from gluing plastic-molds to conventionalsemiconductor-processing techniques (layer deposition, etching, etc.).In this example, the lower portion of the channel 310 is indicated asbeing surface machined into the substrate which in some implementationsincludes a region in and/or on which integrated electronics (such asthose shown in FIG. 2 a) are incorporated for implementation of most orall of the device on a single integrated-circuit chip.

FIG. 4 illustrates yet another biosensor apparatus 400, according to thepresent disclosure, with illustration of a region in the substrate inwhich such integrated electronics are built. The biosensor apparatusincludes a vertically-implemented (nanopore) channel 410 which guidesthe fluid via structures, above and below the substrate surface 415, forproviding the reservoirs 412.

It will be appreciated that the width and length of the channel beingillustrated in FIG. 4 and in the other figures vary depending on theimplementation and/or the bimolecular application. For example, inparticular embodiments, a nanopore channel is implemented with a width(thickness) in the range of about 5 nm-50 nm. Similarly, the channellength can vary widely and in one instance for the width being in therange 5 nm-50 nm, the channel length is about 300 nm. Flow for verysmall pores can be qualitatively different. For instance, for the largerpore sizes (10 s to order or 100 Debeye length diameters), the electrodebias can dramatically change (and hence control) the fluid flow. Asstated herewith in connection with exemplary experimental studiesdocumented in the Appendices A, B, C and D included as part of theabove-noted provisional patent document, these changes can besignificant.

FIG. 5 is an illustration for embodiments, also consistent with theinstant disclosure, showing use of charge tags and antibody/antigenreaction to achieve high specificity and high sensitivity. Asillustrated in FIG. 5 a, the target anti-gen molecule with lowbiological charge translocates through two sensing electrodes andinduces a charge signal in each electrode. The time interval between thetwo correlated signals can be used to deduce its translocation velocity.The signal amplitude is low because of low charge.

In FIG. 5 b, an antigen molecule specific to the target antibody isprepared so that it is cross-linked with a small tag molecule with highcharge (poly-Lysine molecule in this example). This tagged antigenitself alone generates high signal amplitude as well as hightranslocation velocity because of the highly charged tag.

In FIG. 5 c, when the tagged antigen specifically binds to the targetantibody, the immunological complex has high charge and still generateshigh signal amplitude, but the larger size of the complex reduces itstranslocation velocity as compared to FIG. 5 b. This unique, combinedcharge and velocity signature of the immunological complex is thereforeuseful to detect the presence of the target antibody, with its highspecificity coming from the immunological reaction.

For embodiments consistent with the present disclosure, FIGS. 6 a-6 dare graphs (relating to the Appendix and its FIGS. 2( a)-2(d)) useful inshowing that with proper electrical biasing, long-range electrostaticscan be used to create zones that either accumulate or deplete chargedspecies in solutions. For the accumulation case, this approach can beused to concentrate the biomolecule analytes and therefore enhance thesensitivity. For the depletion case, the approach can be used fordifferent applications in desalination, e.g., to remove dissolved saltsfrom sea water or brackish water. In other embodiments and applications,multiple electrodes and different application-specific structures can beused to develop application-specific sensitivities.

The limiting and overlimiting behaviors can be appreciated in connectionwith a specific Id-Vd curve having a constant ΔVg of 1.5 V as in FIG. 6a. For comparison, two more curves are simulated by artificiallyincreasing the solvent viscosity, while keeping the ion mobilityunchanged. The nonlinearity is significantly reduced at higherviscosities, revealing the crucial role of the coupled fluid transport.In the following, four bias conditions are examined specifically fordifferent Vd values, 0.1 V (A), 1.5 V (B), 3 V (C), and 4 V (D), for thecase of nominal viscosity. The four conditions represent differentconductance regimes: linear (A), limiting (B), and overlimiting (C andD).

The limiting and overlimiting conductance is obtained by inspecting thenormalized electrostatic potential, along the nanopore's longitudinalaxis for the four bias conditions, as shown in FIG. 6 b. The effect ofgate potential (ΔVg=1.5 V) is to raise the potential inside thenanopore. This effect is the least significant in the case of low Vd(case A), which is expected because the transport is near equilibriumand the gate potential is strongly screened by counter-ions. Theinsignificant gate modulation is consistent with the linear Id-Vdcharacteristics in this regime. As Vd increases, the counter-ionscreening become suppressed due to enhanced transport, leading tosignificantly increased gate modulation of channel electrostatics. As aresult, in case B there is a highly asymmetric potential profile, wherepotential drop predominantly occurs at one side of the channel. In thisregime, the current is limited by the strong gate modulation. As Vdincreases further so that Vd/2≧ΔVg, the gate potential continues to beless screened but does not exceed its maximum limit ΔVg. Consequently,the relative impact of the gate modulation on the overall channelelectrostatics becomes smaller, as shown by the reduced asymmetry in thenormalized potential (cases C and D). In this regime, the symmetricbiasing condition is reached in the limit as Vd approaches infinity. Theionic conductance therefore asymptotically approaches that under thesymmetric biasing, thus exhibiting the overlimiting behavior.

Concentration Polarization (“CP”) is also observed as a result of theelectrical gating. In FIG. 6 c, the normalized total ion concentration,C=(C+C)/2C, is shown in algorithmic scale for the four bias conditions.In general, the ions deplete and accumulate at the bottom and topportions of the channel, respectively. The magnitude of CP is found tocorrelate strongly with the potential profiles in FIG. 2 b. The CP isinsignificant in case A and becomes stronger in case B. It thengradually reduces as the over limiting regime is reached in cases C andD. For such a correlation, the ion concentration is low in thehigh-field region and vice versa to maintain ionic flux continuity.

The fluid transport is examined in FIG. 6 d, where the solvent velocityfields are shown for the four bias conditions. Induced electro-osmoticflow along the vertical direction is observed with an increasingmagnitude from case A to case B. As the overlimiting regime is reachedin case C, vortex formation is observed in the ion depletion zone at thebottom. The vortex flow is further enhanced in case D, and this trendcontinues as the symmetric biasing condition is asymptoticallyapproached at higher level setting of V_(d). The impact of coupled fluidtransport via certain embodiments in the instant disclosure is tosuppress the current in the limiting regime rather than to raise it inthe overlimiting regime. As indicated via FIG. 6 d, the conductancedifference due to viscosity change is smaller in the overlimiting regimethan in the limiting regime. It is believed that although theirappearances are related, the vortex formation is not the dominant causeof the overlimiting conductance in such a gated nanopore device.

Experimental: Specific Examples and Embodiments

Certain aspects of the present disclosure involve introducing the flowof charged bio-molecules through synthetic pores and detecting inducedpotential signals at the pore walls. As a demonstration, havingsimulated a cylindrically symmetric model system where a piece of 60base dsDNA is located at the center of an aqueous pore. ThePoisson-Nernst-Planck (PNP) equations have been solved to model the iontransport across the pores. Such a continuum-based simulation approachhas been carefully compared and validated with Brownian Dynamicsimulation results for pores with radii >˜2Λ_(D), which is the regime ofinterest here. A general partial differential equation solver, Prophet,has been used to solve those nonlinearly coupled model equations. Aprevious study was based on the same simulation platform to examineenergetic preferred orientations of proteins (mitochondrial creatinekinase, type I hexokinase and cytochrome c) upon immobilization oncharged surfaces. A. H. Talasaz, M. Nemat-Gorgani, Y. Liu, P. Stahl, R.W. Dutton, M. Ronaghi, and R. W. Davis, “Prediction of proteinorientation upon immobilization on biological and nonbiologicalsurfaces,” Proc. Natl. Acad. Sci., 103(40) pp. 14773-8 (2006).Simulation results in this study involve the change of electrostaticpotential due to the presence of the charged biomolecule is plotted. Inequilibrium, Debye-Huckle charge-screening prevents the detection ofpotential changes beyond a few Debye lengths from the chargedbiomolecule. The Debye length at 1 mM KCl concentration is only 10 nm.When an electrical bias is applied to introduce electro-diffusion flow,the screening behavior is qualitatively different: the potential changebecomes long-range and appreciable (˜ mV range) even inside thedielectric membrane. Such a dramatic suppression of the electrostaticscreening is attributed to the current flow in the radial direction. Itcan be analytically derived that in the presence of current flow, along-range, Ohmic-type behavior is superimposed to the Debye-Huckle typebehavior in the solution of charge-induced potential drop. In order toexamine the capability of the proposed device in spatially resolvingcharge profiles, a series of simulations were conducted for molecules,respectively with monopole and dipole (D˜300 electron×nm) chargeprofiles, at different vertical locations in the center of the pore. Thesimulated profiles of potential change sensed in the membrane showsignatures of the different charge profiles. It is also noted that thesignal strength in the Dipole case can be 40× lower than that of theMonopole case, indicating high sensitivity will be needed forapplications of spatially resolving charge profiles.

Further simulations have been carried out to study a realistic devicestructure that is under fabrication for this research project. Inparticular, the sensing metal electrode is sandwiched between two oxideinsulator layers experimental embodiment and accounting for thecircuitry operation of the charge sensing amplifier, the sensingelectrode is connected to the virtual ground. Assuming the biomoleculecharge is −Q, the induced charge in the metal electrode is Q′=βQ, whichis essentially the amount of charge sensed by the amplifier circuitry.With the ratio factor β calculated for different pore radii and biasingconditions, with moderate electrical biasing, the charge that can besensed amounts to ˜50% of the biomolecule charge for a 500 nm poreradius.

By using a foundry CMOS technology for the membrane substrate, the porecan be formed through a complex stack of metal and dielectric layers,which will be used to advantage in manipulating the charged biomoleculein the nanopore. In addition, the charge-sensing amplifier can beintegrated on the same silicon substrate, adjacent to the pore, whichminimizes the charge-signal loss due to parasitic capacitance. The CMOStechnology includes a final deep dielectric-etch using the upper metallayer as a mask. Although this process could be used to etch the porethrough the metallization stack, the minimum size of the pore would be 2μm, which is too large for high-resolution charge detection. Therefore,a focused ion beam (FIB) can be used to etch the nanopore through themetallization stack; pore diameters of 250 nm or even smaller arefeasible. A conventional deep-silicon dry etch or an anisotropic wetetch can be used to etch through the silicon substrate, in order toaccess the nanopore from the backside of the chip.

The low-noise, such as CMOS front-end charge amplifier has an area of240×120 μm² and draws a bias current of 900 μA from the 1.8 V powersupply, for a power consumption of 1.6 mW. Its design is largely basedon front-end charge detectors used for nuclear imaging applications. G.De Geronimo, P. O'Connor, V. Radeka, and B. Yu, “Front-end electronicsfor imaging detectors,” Nuclear Instruments and Methods in PhysicsResearch A 471, pp. 192-199 (2001). The minimum detectable charge can beset by the amplifier's input-referred noise or by the noise current dueto the ionic current through the nanopore. Since the amplifier is notconstrained in its power dissipation for this application, withinreasonable limits, the lowest equivalent noise charge (ENC) can beachieved by matching its input capacitance to be equal to thesense-electrode and parasitic capacitances. For an example nanopore,this capacitance is approximately 70 fF. Assuming a typical noise modelfor the amplifier, the input-referred noise voltage has aroot-mean-square value of around 10 μV, over a measurement bandwidth of1 kHz to 1 MHz, which translates into an ENC=9 electrons.

The presence of a DC ionic current through the nanopore is also a sourceof noise, which must be considered in estimating the minimum detectablecharge. The impedance field method can be used to estimate the ioniccurrent-induced charge noise at the sense electrode, using a techniquesimilar to that used to model the gate-induced noise current in MOStransistors. F. Danneville, H. Happy, G. Dambrine, J.-M. Belquin, and A.Cappy, “Microscopic noise modeling and macroscopic noise models: howgood a connection?,” IEEE Trans. On Electron Devices, vol. 41, no. 5,pp. 779-786 (May 1994). Current fluctuations across the pore couple tothe sense electrode and contribute to an equivalent input-referredcharge noise, when integrated over the measurement bandwidth. Theestimated ENC from this source is 25 electrons, which dominates theamplifier's input-referred noise. This result is very encouraging, sinceit indicates that single-stranded DNA can be detected with about 15 basepairs (30 electrons).

The multiple electrodes in the walls of the nanopore allow theelectrostatic manipulation of the charged biomolecule's translocationthrough the pore. By using lock-in techniques (such asphase-locked/frequency-locked loop circuits), enhanced sensitivity canbe achieved by limiting the measurement bandwidth.

According to another important aspect of the disclosure, integratedsensors are developed for direct detection of single target molecules byleveraging a nanotechnology-enabled type of sensing mechanism forquantifying the net charge of biomolecules passing through a nanoscaleconfined geometry. Here, a first characterization of the electricalproperty of the nanopore front-end and then the integrated platform.After the optimization of the sensor for an intended application, use ofthe device for detection of target DNAs hybridized to the specificprobes. Finally, use of the multi-electrode actuation strategies forincreasing the detection limit of the device with the final goal ofcounting the DNA bases with single base resolution.

A first more specific example involving the fabrication andcharacterization of charge sensitive nanopore devices: characterizationof MEMS nanopore devices that are already fabricated without integratedelectrical circuitry; and integration of the electrical circuitry withthe nanopore device to increase the detection sensitivity.

The development of the herein described charge sensor embodimentsinvolves coherent efforts in design, fabrication and characterization atboth device and circuit levels to ensure that the front-end nanoporesand the charge sensing circuits are properly designed and operational.Owing to the multi-layer electrode fabrication capabilities offered bythe advanced IC technologies, a very large design space is expected tobe explored to achieve optimal performance. Extensive characterizationand testing on those already fabricated, prototypical MEMS devices aswell as those integrated-circuitry devices under fabrication arenecessary to provide important design parameters such as the noise floorand parasitic capacitances. In particular, the noise performance of thecharge sensitive amplifier is dependent on front-end parasiticcapacitance. Subsequent efforts to system performance optimization willbe based on those design parameters.

Prototypical front-end nanopore devices (without integrated circuits)are fabricated using MEMS fabrication process. FIB techniques are usedto create pores in those devices. SEM imaging is used to examine thesize and shape of the pores, thus giving useful information on the rangeof pore sizes that are manufacturable with FIB. The electrical testingof such devices is then performed by connecting external AC/noiseanalyzers to the sensing electrode through wire bonding. Biosensordevices with integrated amplifier circuit are fabricated using advancedIC process. In particular, various design schemes can be implemented byutilizing the multi-layer electrode fabrication capabilities. An actuallay-out design of integrated sensors can use 0.18 μm CMOS technologywith a low noise amplifier occupying an area of 240 um×120 um with apower consumption of 1.6 mW included and fabricated as part of the 0.18μm CMOS process. As an example experiment, the primary limit on theirproximity to the pores allows at least a 100 um electronics exclusionradius for the backside etch. The electrode designs range from a singleelectrode to a three-layer 6-electrode stack. Multiple electrode stacksallow enhanced time domain resolution as well as electrostatic controlof the charged target molecule. After the fabrication process at thefoundry facilities, these nanopore test chips have a backside cavityetch (200 μm diameter) to remove the substrate silicon around the poresites. Then the nanopores are precision etched by FIB from thefront-side, through the CMOS metal stack. After the FIB post-processingstep, the nanopore sensor fabrication is complete, with optional testingto characterize parasitic capacitance, noise floor including theelectronic noise component and sensitivity of the charge-sensingamplifier circuit.

For targeted minimum detectable signal of the integrated sensors at anarbitrarily-selected level of sub-100 electrons, a significant area ofthe fabricated sensor chip may need to be immersed in the ionicsolution. For example, a poorly-selected/implemented O-ring seal used ina nanopore implementation may result in large parasitic capacitance andsignificant coupling of the ionic solution noise. Solutions include (i)using IC processes to increase the oxide layer thickness to reduceparasitic capacitance; and (ii) introduction of ionic solution usinghypodermic needles which have much reduced contact area. Also, becausethe system minimum detectable signal is limited by the noise,particularly flicker noise, in the electronic amplifier circuit, buriedchannel MOSFETs are used as active elements in the amplified design. Theburied channel MOSFETs are known to significantly reduce the flickernoise, as shown in previous extracellular action potential study. S.Ingebandt, “Characterization of the cell-transistor coupling,”dissertation thesis, Johannes Gutenberg University Mainz, Germany(2001).

A second specific example involves the detection of the target DNA bycharge nanopore devices. To develop an electrical sensor for label-freedetection of single target DNA molecules, the proposed platform can beused for many genomic biomarker detection applications, such as viraland bacterial detection, without any PCR amplification step. This can beachieved by incubating the target DNAs with their specific probeoligonucleotides and employing nanopore charge sensing/sensor to detectthe hybridization event by counting the net charge of each translocatingDNA molecules. The nanopore charge sensor can detect the translocationevents of the DNA molecules passing through the pore, with a train ofpulses at the output of the integrated charge sensor. Then the chargesensor is used to distinguish ssDNA molecules from dsDNA. Initially, theelectrical signals due to the translocation of a 1 kb PCR product arerecorded and compared to the signals from the 1 kb ssDNA. Pulses werethen realized with twice the amplitude for double-stranded PCR products.Subsequently, a gradual reduction of the length of DNA strands tomeasure the minimum detection limit of the device. Based on thisinformation, a specific probe is designed so that the charge sensor candetect its hybridization to the target DNA.

When such a device is used for analyzing DNA molecules, there is anon-desirable interaction between translocating DNA and the electrode inthe nanopore and possible electrochemical reactions. Alternativepost-processing steps are implemented to passivate the sensing electrodethat is exposed to the solution. Such surface treatments may involvedeposition of metal oxide insulating layers by atomic layer depositionor surface chemistry treatment. The electrode passivation also minimizesFaradic current component across the interface. Also, if the detectionlimit of the device is not sufficient, the salt concentration is loweredin the solution to further reduce the charge screening effect andincrease the induced charge.

In a third specific example, measuring of the DNA length by counting theDNA charges with the final goal of single base resolution and thenevaluating the resolution of the charge sensor in quantification of thenumber of DNA bases. The sensitivity of the sensor is enhanced bymultiple readout of the charge by multiple electrodes in the nanopore;electrical focusing of the DNA to the nanopore wall; andelectrical-control of the DNA translocation speed by using multipleelectrodes in the pore.

The ability to measure DNA charge to single electron precision resultsin quantification of DNA length to single base level, which enablesimportant biological and medical applications. This challenging aspectis an important achievement of the sensor.

In one example approach, this aspect is implemented with aggressivesignal detection and processing to dramatically improve SNR, e.g., byusing any of various detection schemes. As examples: (a) Multiplexing:multiple metal layers used as independent sensing electrodes, eachconnected to its own amplifier and analog-to-digital converter (ADC).The multiple signals are multiplexed and processed in a digital signalprocessing (DSP) unit. The time delayed signals can be correlated toprovide enhanced signal. (b) Deceleration: auxiliary electrodes are usedto create a drift region in one segment of the aqueous pores, where theelectrical field is small by design. For nanopores without such driftingdesigns, the DNA translocation time is usually in the sub-millisecondrange. The purpose of this example design is to increase the dwellingtime of DNAs in the drift region where the sensing electrode is, so thatthe actual signal sampling is enhanced. (c) Steering: with thesuppressed screening effect, electrical field is introduced in thetransverse planes so that, with the suppressed screening effect, thecharged biomolecules can be electrically steered closer to the surfaceof the sensing electrode for improved signals.

As an alternative to enhance the above signal detection schemes andfacilitate achieving resolution down to the signal electron level,AC-based signal detection is introduced to further boost sensitivity.Using this approach: (a) the charge sensing can be moved away from thelow-frequency regime where flicker noise usually compromises signaldetection; and (b) advanced lock-in amplification techniques can beemployed, which have been broadly used in instrumentations to extractsignals for SNR, for the experimental embodiments, as low as −60 dB.

By correlating signals from two or more sense electrodes, thebio-molecule velocity can be measured, i.e., besides the chargeinformation, to also obtain the size information simultaneously. Thatmeans multiple parameters are available to identify the target.Experimental work has been focused on nanopore bio-sensors using theelectro-diffusion effect. A new device fabrication process withprotective Atomic Layer Deposition coatings on 500 nm size,photo-lithographically-defined pores can be used to improve the yieldsignificantly, e.g., up to 80%. Using custom charge sense amplifiercircuits or commercially-available circuits, significant biologicalsensing advancements can be realized. Also, for testing the operationprinciples, silicon nanowire prototypes are believed to be useful.

In accordance with yet further aspects, it has been discovered herewiththat the movement of bio-molecules and ionic species can be controlledon this platform, i.e. nano-channels with side electrodes, based on thesame or similar long-range electrostatic effect.

Based on other practical applications, the present disclosure is alsodirected to the following embodiments (including uses) based upon andleveraging from many of the previously discussed aspects, alone and/orin combination. First, use of a nanostructure as described above is notnecessarily to be functionalized by a molecule type. In this case, thesensing mechanism is used to detect the charges of biomolecules flowingby the nanostructure rather than of those attached to the nanostructure.An example is the use of silicon nanowires or carbon nanotubes toreplace the sensing metal electrodes (e.g., 320 in FIG. 3). As comparedto the metal electrodes, the silicon nanowires or carbon nanotubes havetheir intrinsic trans-impedance gain at the very front end, which helpsto minimize parasitic and offers better sensitivity.

As another aspect, by correlating signals from two or more sensingelectrodes or nanostructures, the bio-molecule translocation velocitycan be deduced in addition to its charge. The combined signature ofvelocity and charge provides high-quality identification ofbio-molecules as compared to charge signal alone. One example of usingthis combined signature is described in FIG. 5. Use of the nanochannelswith control electrode arrays, as illustrated in connection with theexample of FIG. 2 a, to electrically modulate the concentration of ionsor charged molecules. The channel size is not limited by the Debyescreening length as set forth herein.

Yet another aspect involves electrical biasing of the control electrodearrays to create solution zones between the electrodes with eitherdepleted or accumulated concentrations of ions or charged moleculesdepending on the biasing schemes. The depletion and accumulation modesof operation can be used together and also used disparately in entirelydifferent domains.

Other embodiments use electrical biasing in different manners. In onesuch embodiment, two control electrodes are used to create a solutionzone between them, where the electrical field along the channeldirection is high. In this zone, the ion concentration is depletedbecause of the high electrical field in order to maintain ionic fluxcontinuity. The application of this depletion mode operation includeswater desalination or de-ionization.

In another embodiment of this type, electrical biasing of two controlelectrodes is used to create a solution zone between them, where theelectrical field along the channel direction is low. In this zone, theion concentration is accumulated because of the low electrical field inorder to maintain ionic flux continuity. Applications of thisaccumulation mode operation include biological sample pre-concentration.

In yet other embodiments, side channels are connected to the depletionor accumulation zones to extract out the solution of depleted oraccumulated concentrations of ions or charged molecules. The extractioncan be achieved using either pressure or voltage difference.

These above-described embodiments can be used alone or in variouscombinations.

Each of the Appendices (A, B, C and D) of the above-noted ProvisionalPatent document includes further experimental embodiments and featuresdescribed and illustrated in connection with the above-describedembodiments. These embodiments and features can be implemented asdescribed or, as would be recognized, aspects thereof can be used in avariety of combinations. For further background information regardingcharge detection of bio-molecular substances and/or affinity-typeapproaches to bio-molecule detection, reference may be made to thereferences listed in each of the attached Appendices A, B, C and D. Forexample, U.S. Pat. No. 6,413,792 (Ultra-Fast Nucleic Acid SequencingDevice and a Method for Making and Using the Same) and U.S. Pat. No.7,001,792 (Ultra-Fast Nucleic Acid Sequencing Device and a Method forMaking and Using the Same) provide ample discussion of previousstructures and methods in specific DNA-directed applications. Thearticle, Effect of Electrodiffusion Current Flow on ElectrostaticScreening in Aqueous Pores, J. Appl. Phys. 103, pp. 084701-1-03 (2008),provides a mathematical discussion of the Debye-Huckel theory and chargesensing being limited to within a few Debye lengths from thebiomolecules. Each of these references is fully incorporated herein byreference, generally and specifically.

Further, the Appendix attached hereto (forming part of the instantpatent document) and entitled, Limiting and overlimiting conductance infield-effect gated nanopores, forms part of this patent document andalso is incorporated by reference as further describing the above andrelated embodiments, uses and applications by way of its recentpublication bearing the same title: Y. Liu, D. Huber, and R. Dutton,Appl. Phys. Lett. 96, 253108 (2010).

While the present invention has been described with reference to severalparticular example embodiments, those skilled in the art will recognizethat many changes may be made thereto without departing from the spiritand scope thereof. The present invention is applicable to a variety ofsensor implementations and other subject matter, in addition to thosediscussed above.

What is claimed is:
 1. A method for detecting bio-molecules, the methodcomprising: providing a channel having an effective width that is notlimited by the Debye screening length; containing and presenting asample having bio-molecules for delivery in the channel; using a firstelectrode and a second electrode, electrically coupling a charge in thesample to enhance ionic current flow therein; using a sense electrodelocated along the channel, sensing a characteristic of the biologicalsample at least partly based on electrostatic interaction between theenhanced ionic current flow of the sample and the sense electrode; andprocessing a charge signal sensed and carried from the sense electrodeand providing an output indicative of a signature for the bio-moleculesdelivered in the channel.
 2. The method of claim 1, wherein using afirst electrode and a second electrode including providing a modulationsignal and further using lock-in detection technique to lock onto themodulation signal.
 3. A circuit-based arrangement comprising: asubstrate securing a channel located along a surface of the substrate,the channel having an effective width that is not limited by the Debyescreening length; at least one reservoir configured for containing andpresenting a sample having bio-molecules for delivery in the channel; afirst electrode and a second electrode, each of the electrodesconfigured for electrically coupling a charge in the sample to enhanceionic current flow therein; at least one sense electrode located alongthe channel and configured for sensing a characteristic of thebiological sample at least partly based on electrostatic interactionbetween the enhanced ionic current flow of the sample and said at leastone sense electrode; and a charge-signal processing circuit coupled tosaid at least one sense electrode for processing a signal carriedtherefrom and providing an output indicative of a signature for thebio-molecules delivered in the channel.
 4. The circuit-based arrangementof claim 3, wherein said at least one reservoir includes a firstreservoir and a second reservoir, each of the first and secondreservoirs fluidly-coupled to the channel.
 5. The circuit-basedarrangement of claim 4, wherein the first electrode and the secondelectrode are respectively located in the first reservoir and in thesecond reservoir.
 6. The circuit-based arrangement of claim 3, whereinthe channel has an effective width that is greater than or equal toabout 200 nm.
 7. The circuit-based arrangement of claim 3, wherein atleast one of the electrodes is buried in the substrate adjacent thechannel.
 8. The circuit-based arrangement of claim 3, wherein said atleast one of the electrodes is buried in the substrate adjacent thechannel and arranged for charge sensing.
 9. The circuit-basedarrangement of claim 3, wherein said at least one of the electrodes isburied in the substrate adjacent the channel and arranged forbio-molecule movement control.
 10. The circuit-based arrangement ofclaim 3, wherein said at least one of the electrodes includes a pair ofelectrodes buried in the substrate adjacent the channel and arranged forcharge sensing and bio-molecule movement control via a modulationtechnique, and wherein the channel is a surface channel arrangedlaterally.
 11. The circuit-based arrangement of claim 3, furtherincluding nanostructures disposed along the channel for sensing chargesof the bio-molecules.
 12. The circuit-based arrangement of claim 3,wherein at least one sense electrode is configured for providing ACmodulation and a lock-in technique, and for sensing bio-molecule chargesat the single-molecule level.
 13. The circuit-based arrangement of claim3, wherein at least one sense electrodes is configured using a pluralityof sense electrodes, each driving an input of a sense amplifier.
 14. Thecircuit-based arrangement of claim 3, wherein charge sensing is notlimited by electrolyte screening.
 15. The circuit-based arrangement ofclaim 14, wherein at least one sense electrode and the charge-signalprocessing circuit are configured for providing the sensing ofbio-molecule charges at the single-molecule level.
 16. The circuit-basedarrangement of claim 3, further including a modulation circuitconfigured to provide a modulation signal via the first and secondelectrodes, and wherein the charge-signal processing circuit includes alock-in detection circuit configured to lock onto the modulation signal.17. The circuit-based arrangement of claim 3, wherein at least one senseelectrode is located along the channel concurrently with the biologicalsample in the channel, whereby the first and second electrodes sense theelectrostatic interaction.
 18. An apparatus comprising: a slab securinga channel, the channel having an effective width that is not limited bythe Debye screening length; at least one reservoir configured forcontaining and presenting a sample having bio-molecules for delivery inthe channel; a first electrode and a second electrode, each of theelectrodes configured for electrically coupling a charge in the sampleto enhance ionic current flow therein; at least one sense electrodelocated along the channel and configured for sensing a characteristic ofthe biological sample at least partly based on electrostatic interactionbetween the enhanced ionic current flow of the sample and said at leastone sense electrode; and a charge-signal processing circuit coupled tosaid at least one sense electrode for processing a signal carriedtherefrom and providing an output indicative of a signature for thebio-molecules delivered in the channel.
 19. The apparatus of claim 18,wherein the slab is configured to secure the channel along a surface ofthe slab.
 20. The apparatus of claim 18, wherein the slab includes asubstrate that is configured to secure the channel along a surface ofthe substrate.
 21. The apparatus of claim 18, further including a guidemember attached to a surface of the slab and configured to provide thechannel.
 22. The apparatus of claim 18, wherein the slab is configuredto secure the channel vertically using through-holes.
 23. The apparatusof claim 18, wherein the sense electrode includes a nanostructure thatis functionalized with a molecule type and conductance characteristic ofthe nanostructure in response to a target type of bio-molecule in thechannel.
 24. An apparatus comprising: channel means for providing achannel having an effective width that is not limited by the Debyescreening length; reservoir means for containing and presenting a samplehaving bio-molecules for delivery in the channel; means, including afirst electrode and a second electrode, for electrically coupling acharge in the sample to enhance ionic current flow therein; sensingmeans, including a sense electrode located along the channel, forsensing a characteristic of the biological sample at least partly basedon electrostatic interaction between the enhanced ionic current flow ofthe sample and the sense electrode; and a charge-signal processingcircuit coupled to the sensing means and configured for processing asignal carried therefrom and providing an output indicative of asignature for the bio-molecules delivered in the channel.